U.S. patent number 8,992,967 [Application Number 14/062,442] was granted by the patent office on 2015-03-31 for poly (diol-co-citrate) hydroxyapatite composite for tissue engineering and orthopaedic fixation devices.
This patent grant is currently assigned to Northwestern University. The grantee listed for this patent is Northwestern University. Invention is credited to Guillermo Ameer, Hongjin Qiu, Jian Yang.
United States Patent |
8,992,967 |
Ameer , et al. |
March 31, 2015 |
Poly (diol-co-citrate) hydroxyapatite composite for tissue
engineering and orthopaedic fixation devices
Abstract
The present invention is directed to a novel poly(diol
citrates)-based bioceramic composite materials created using
completely biodegradable and a bioceramic material polymers that
may be used in implantable devices. More specifically, the
specification describes methods and compositions for making and
using bioceramic composites comprised of citric acid copolymers and
a bioceramic material.
Inventors: |
Ameer; Guillermo (Chicago,
IL), Qiu; Hongjin (Evanston, IL), Yang; Jian
(Arlington, TX) |
Applicant: |
Name |
City |
State |
Country |
Type |
Northwestern University |
Evanston |
IL |
US |
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Assignee: |
Northwestern University
(Evanston, IL)
|
Family
ID: |
38345803 |
Appl.
No.: |
14/062,442 |
Filed: |
October 24, 2013 |
Prior Publication Data
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Document
Identifier |
Publication Date |
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US 20140155516 A1 |
Jun 5, 2014 |
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Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
Issue Date |
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13309014 |
Dec 1, 2011 |
8568765 |
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11704074 |
Feb 8, 2007 |
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60771241 |
Feb 8, 2006 |
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Current U.S.
Class: |
424/426; 424/422;
424/423 |
Current CPC
Class: |
A61L
27/12 (20130101); A61L 31/123 (20130101); A61L
27/18 (20130101); A61L 27/46 (20130101); A61L
31/148 (20130101); A61L 31/127 (20130101); A61L
31/121 (20130101); A61L 27/427 (20130101); A61L
31/06 (20130101); A61L 27/425 (20130101); A61L
31/026 (20130101); A61L 27/58 (20130101); A61L
2430/02 (20130101) |
Current International
Class: |
A61F
2/00 (20060101) |
References Cited
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|
Primary Examiner: Rogers; James
Attorney, Agent or Firm: Casmir Jones, S.C.
Parent Case Text
CROSS REFERENCE TO RELATED APPLICATIONS
The present application is a continuation of pending U.S. patent
application Ser. No. 13/309,014 filed Dec. 1, 2011, which is a
continuation of abandoned U.S. patent application Ser. No.
11/704,074, filed Feb. 8, 2007, which claims benefit of U.S.
Provisional Application No. 60/771,241 filed Feb. 8, 2006, each of
which is incorporated herein by reference in its entirety.
Claims
We claim:
1. A composition comprising a composite of: a) a polyester
comprising one or more linear aliphatic diol monomers and a citric
acid monomer; and b) a bioceramic used for implantable tissue
devices, wherein said composition comprises said bioceramic in an
amount of 55 wt % or greater.
2. The composition of claim 1, wherein the linear aliphatic diol
monomer comprises between about 2 and about 20 carbons.
3. The composition of claim 2, wherein the polyester comprises
repeating units of the same linear aliphatic diol monomer.
4. The composition of claim 3, wherein the linear aliphatic diol is
1,8-octanediol.
5. The composition of claim 4, wherein the polyester is poly
1,8-octanediol co-citric acid.
6. The composition of claim 2, wherein the polyester is poly 1,10
decanediol co-citric acid.
7. The composition of claim 2, wherein the polyester comprises
repeating units of different linear aliphatic diol monomers.
8. The composition of claim 1, wherein said bioceramic is selected
from the group consisting of calcium phosphate bioceramics,
aluminabased bioceramics; zirconia-based bioceramics; silica-based
bioceramics, and pyrolytic carbon-based bioceramics.
9. The composition of claim 8, wherein said bioceramic is a calcium
phosphate bioceramic.
10. The composition of claim 8, wherein said bioceramic is
hydroxyapatite (HA).
11. The composition of claim 8, wherein said bioceramic is a
tricalcium phosphate.
12. The composition of claim 1, wherein said bioceramic is present
in an amount of 55 wt % to 95 wt %.
13. The composition of claim 1, wherein said bioceramic is present
in an amount of 65 wt % to 95 wt %.
14. The composition of claim 1, wherein said bioceramic is present
in an amount of 70 wt % to 95 wt %.
15. An implantable device comprising the composition of claim
1.
16. The composition of claim 1, wherein the bioceramic comprises
particles.
17. The composition of claim 16, wherein the particles have a size
from a nanometer to a micron.
Description
FIELD OF THE INVENTION
The present invention describes new composites for use in
orthopedic devices.
BACKGROUND
Orthopaedic, cranio-facial, and oral-maxillofacial surgeons often
use tissue fixation devices such as pins, plates, and screws that
are made from poly-1-lactide (PLLA), a biodegradable polymer[1-7].
Although biodegradable devices can have significant advantages over
their metal counterparts, there are concerns with their use. These
include slow degradation, which can be as long as 5 years, and
their inability to fully integrate with bone, which can be a
problem for revision surgeries [8, 9]. Also, PLLA bone screws can
fracture during the fixation procedure. A strategy to improve the
osteointegration capacity of PLLA has been to blend it with
hydroxyapatite (HA), a bioceramic that can be found in natural bone
mineral. Although HA is very brittle and hard to process into
fixation devices of sufficient strength and fatigue resistance, it
can impart osteoconductivity to polymers [10, 11]. Researchers have
shown that under certain conditions, addition of HA particles can
improve the mechanical properties of the polymer component when
used in a composite blend [12-14]. Therefore composites of polymers
with bioceramics may be a suitable compromise to meet mechanical
property requirements and achieve osteointegration of the implant.
Nevertheless, there remains a significant problem in that the PLLA
continues to slowly degrade over a period of time. Moreover,
incorporation of more than about 30 wt. % of HA into any such
composite leads to a material that is too brittle for use in
implantable devices. Thus, there remains a need for composite
materials that are biocompatible, can be easily processed and will
fully integrate with the surrounding bone and tissue within a year
of implantation.
SUMMARY OF THE INVENTION
The present invention provides a composition comprising a composite
of a citric acid polyester having the generic formula (A-B-C)n,
wherein A is a linear aliphatic dihydroxy monomer; B is citric
acid, C is a linear aliphatic dihydroxy monomer, and n is an
integer greater than 1; and a bioceramic used for implantable
tissue devices, wherein less than 75 wt. % weight ratio of said
composition comprises said bioceramic. In other embodiments, at
least 30 wt. % weight ratio of said composition comprises said
bioceramic. In specific and alternative embodiments, the bioceramic
component forms about 5 wt. %, 10 wt. %, 15 wt. %, 20 wt. %, 25 wt.
%, 30 wt. %, 35 wt. %, 40 wt. %, 45 wt. %, 50 wt. %, 55 wt. %, 60
wt. %, 65 wt. %, 70 wt. %, 75 wt. %, 80 wt. %, 85 wt. %, 90 wt. %,
95 wt. %, or greater than 95 wt. % of the composition.
Preferred compositions are those in which A is a linear diol
comprising between about 2 and about 20 carbons. In other preferred
compositions, C is a linear diol comprising between about 2 and
about 20 carbons. In still other preferred compositions, both A and
C are the same linear diol. In alternative embodiments, A and C are
different linear diols. In specific compositions, the linear diol
is 1,8-octanediol.
In exemplary embodiments, the linear aliphatic dihydroxy poly
1,8-octanediol co-citric acid. In still other exemplary
embodiments, the linear aliphatic dihydroxy poly 1,10-decanediol
co-citric acid.
The bioceramic may be any ceramic typically used in medical
applications. For example, exemplary such bioceramics may be
selected from the group consisting of calcium phosphate
bioceramics, alumina-based bioceramics; zirconia-based bioceramics;
silica-based bioceramics, and pyrolytic carbon-based bioceramics.
Combinations of bioceramics may be used. In certain preferred
embodiments, the bioceramic is a calcium phosphate bioceramic at a
weight percentage of from 30 wt. % to about 75 wt. % of the total
weight of the composition. In other embodiments, the bioceramic is
hydroxyapatite (HA) at a weight percentage of between about 40 wt.
%.+-.5 wt. % to about 70 wt. %.+-.5 wt. % HA to 35 wt. %.+-.5 wt. %
to about 25 wt. %.+-.5 wt. % citric acid polyester.
The compositions of the invention are such that they produce
composites of preferred bending strength. In preferred embodiments,
the composite has a bending strength of from about 33.9 to about
41.4 MPa. In still other preferred embodiments, the composites have
a preferred compression strength, wherein the compression strength
is preferably from about 32 to about 75 MPa. The composites
alternatively or in addition may have a defined tensile strength in
the range of from about 6 to about 10 MPa. The composites also may
be characterized according to their shear strength. The preferred
shear strength of the composites is between about 23 to about 28
MPa. The composites may be defined according to their bending
modulus. The bending modulus is preferably from about 0.275 to
about 0.502 GPa. In other embodiments, the composites are
characterized by having a compression modulus of from about 0.19 to
about 0.45 GPa. In still other embodiments, the composites may be
characterized by a tensile modulus of from about 0.02 to about 0.34
GPa. The composites of the invention may advantageously be
characterized by one or more of these features.
In addition, the invention contemplates composites of the invention
which in addition to the citric acid polyester and bioceramic
further comprise a polymer selected from the group consisting of
poly(hydroxyvalerate), poly(lactide-co-glycolide),
poly(hydroxybutyrate), poly(hydroxybutyrate-co-valerate),
polyorthoester, polyanhydride, poly(glycolic acid),
poly(glycolide), poly(L-lactic acid), poly(L-lactide),
poly(D,L-lactic acid), poly(D,L-lactide), poly(caprolactone),
poly(trimethylene carbonate), and polyester amide.
In specific embodiments, the composition is molded into an
orthopedic fixation device. Preferred such devices include but are
not limited to bone screws, bone pins, bone rods, and bone
plates.
Also contemplated herein is an artificial bone, wherein said bone
is comprised of a composite of the present invention.
Another aspect of the invention describes a substrate for use in an
implantable device comprising a composite of the invention
prepared, molded or fabricated into an orthopedic fixation device
or an artificial bone structure. In preferred embodiments, the
substrate further comprises a surface modification to facilitate
implantation of said device with a decreased risk of implant
rejection.
Also provided herein is a method of producing an implantable
device, comprising: preparing a composition according to the
present invention and molding the composition into an orthopedic
fixation device or an artificial bone for implantation. Such a
method may be used in the preparation of a bone screw, a bone pin,
and a bone plate.
Also provided herein is an implantable device comprising a polymer
composition of the present invention.
Other features and advantages of the invention will become apparent
from the following detailed description. It should be understood,
however, that the detailed description and the specific examples,
while indicating preferred embodiments of the invention, are given
by way of illustration only, because various changes and
modifications within the spirit and scope of the invention will
become apparent to those skilled in the art from this detailed
description.
BRIEF DESCRIPTION OF THE DRAWINGS
The following drawings form part of the present specification and
are included to further illustrate aspects of the present
invention. The invention may be better understood by reference to
the drawings in combination with the detailed description of the
specific embodiments presented herein.
FIG. 1 Production of bone screws of POC-HA composite with 40 wt. %
HA obtained by compression molding and machining methods. POC-HA
composites were prepared by an in situ post-polymerization of HA
and pre-POC blending at 80.degree. C. for 3 days and 120.degree. C.
for 1 day under vacuum.
FIG. 2 SEM images of surface of POC-HA composites with a) 40 wt. %
HA and b) 65 wt. % HA.
FIG. 3 Weight loss of POC-HA composites with HA fraction of 40, 50,
60, to 65 wt. % in vitro (PBS at 37.degree. C.) at 2, 6, 12 and 20
weeks. POC-HA composites were prepared at 80.degree. C. for 3 days
and 120.degree. C. under vacuum for 1 day
FIG. 4 Mineralization in SBF for POC at a) 3 days and b) 15 days,
for POC-HA with 40 wt. % HA at c) 3 days, d) 15 days, and for
POC-HA with 65 wt. % HA e) 3 days and f) 15 days. Magnification of
all images: .times.4.5K.
FIG. 5 LM images of POC seeded with HOB in vitro at a) 3 days b) 14
days; SEM images of POC-HA composites seeded with human osteoblasts
in vitro for 40 wt. % HA at c) 3 days d) 15 days and for 65 wt. %
at e) 3 days and f) 14 days.
FIG. 6 is a schematic representation of the synthesis of
poly(1,8-octanediol-co-citric acid).
DESCRIPTION OF THE PREFERRED EMBODIMENTS
In co-pending applications 60/721,687 and PCT/US2004/030631, (each
incorporated herein by reference in its entirety) there is
described the synthesis and characterization of elastomeric and
biodegradable polyesters, referred to as poly(diol citrates (POC);
see PCT/US2004/030631) and composites comprising POC with a second
polymer (see U.S. 60/721,687 and applications depending therefrom).
The mechanical properties and degradation rates of POC polymers can
be controlled with synthesis conditions of the polycondensation
reaction and choice of diol, and their preparation does not involve
any harsh solvents or exogenous catalysts [15, 16]. Furthermore,
poly(diol citrates) can be very inexpensive, relative to
poly(.alpha.-hydroxy acid) biodegradable polymers.
In the present invention, it is shown that a composite of a
poly(diol citrate) with HA would have the desired characteristics
of a bioceramic with improved processability, mechanical
properties, and degradation characteristics.
Poly(1,8-octanediol-co-citrate) (POC; the production of which is
shown in FIG. 6) was selected for the preparation of exemplary
composites of the present invention because of its faster
degradation rate (a few months to 1 year) than PLLA (3-5 years) and
because its mechanical properties can be tailored by simply
changing reaction conditions such as reaction temperature and time,
and the ratio of 1,8-octanediol to citric acid [15, 16]. POC has
also been shown to be biocompatible and could potentially enhance
the biointegration of the surrounding soft tissue as in the case of
fixation of a ligament graft [15, 16]. Moreover, these materials
are inexpensive and easy to synthesize, an additional advantage for
clinical application.
In preferred embodiments composites are prepared from POC. The
methodology that was used to post-polymerize the materials is
unique. Polycondensation of POC can be conducted under no vacuum,
no catalyst, and low reaction temperature (under 100.degree. C.,
such as 60.degree. C., 80.degree. C., even as low as 37.degree.
C.). Catalyst and high temperature could also be applied if
needed.
The present invention shows that the POC-HA composites have the
desired mechanical, degradation, mineralization, and cell
compatibility characteristics to serve as compositions for
implantable devices such as orthopedic fixation devices as well as
to serve as compositions for use in the production of artificial
bone structures. The feasibility of fabricating (i.e., molding, or
machining) composite bone screws of POC-HA by compression molding
and machining also is shown herein.
Compositions of poly(diol citrates) comprise a citric acid
polyester having the generic formula (A-B-C)n, wherein A and C
could be any of the diols or any combination of the diols; B could
be citric acid, malic acid or their combinations. The diols include
aliphatic diols, branched diol, cyclodiol, triol, heteroatom
containing diol (such as N-methyldiethanolamine, MDEA) and
macrodiol or their combinations. Any composites composed of any
biodegradable elastomers (e.g. poly diol-citric acid,
polyurethanes, polycaprolactone and copolymers thereof) and any
bioceramic (such as HA, TCP, OCP and bioglass) with fraction from 0
to 100 wt. %, for example, 30 to 95 wt. % ceramics can be applied
to fabricate fixation devices (such as pins, wires, tacks, plates,
rods, screws) for clinic and cell scaffolds for tissue engineering
and drug delivery.
Orthopedic fixation devices such as bone screws are often used in
orthopedic, craniofacial and oral-maxillofacial surgery. The vast
majority of these devices are made from metals, which can cause
unwanted tissue reactions, and lead to significant bone removal if
a secondary intervention is required. Alternatively, biodegradable
polymers such as poly(L-lactide) (PLLA) have been used for the
fabrication of some fixation devices where significant weight
bearing is not an issue for the proper function of the device.
Unfortunately, these devices, in particular PLLA bone screws, are
not osteoconductive, have a slow degradation rate (3-5 years), can
fracture during the fixation procedure, and are significantly more
expensive than their metal counterparts.
One way to deal with the osteointegration deficiencies of polymers
has been to blend them with bioceramics such as hydroxyapatite (HA)
and tricalcium phosphate (TCP)[12-14, 22]. Bioceramics have been
shown to be osteoconductive, but are brittle and hard to process
into useful fixation devices for orthopaedic applications. As a
result, several researchers have developed and investigated
composites of HA or TCP with poly (a hydroxy acids) such as PLLA.
Studies have found such composites to osteointegrate more readily
than the pure polymer, supporting the further study of HA/polymer
composites. Nevertheless, the polymer component remains a
relatively large percentage of the composites, typically 70 wt. %,
and in the case of PLLA that is used commercially, the time to
total degradation after its function has been completed is still
too long.
As shown in the Examples herein below, the novel bioceramic
composites, based on poly(diol citrates), have enhanced
osteointegration potential relative to current biodegradable
fixation devices. In the present invention it is shown that it is
possible to prepare bone screws and other orthopedic fixation
devices that consist mostly of the bioceramic component. This
maximizes the osteointegration while employing a degradable and
relatively inexpensive elastomer as the macrophase binder. The
inventors showed that poly(1,8-octanediol co-citrate) (POC),
improves the processability and mechanical properties of bioceramic
bone screws due to its biocompatibility, mechanical properties,
controllable degradation rates (a few months to 1 year), and mild
synthesis conditions[15]. An important criterion for the POC-HA
composites was the ability to process samples via molding and
machining methods. The Examples below demonstrate the successful
synthesis of POC-HA composites with HA compositions of 40, 50, 60,
and 65 wt. % that were readily molded and machined to make bone
screws. While HA percentages of 70 or higher may be fabricated,
those composites are not readily amenable to molding or machining
and hence are less desirable, but may nonetheless be useful
composites if molding is not required. An HA content below 40 wt. %
resulted in composites that were too rubber-like and difficult to
machine. Such composites with an HA content of less than 40 wt. %
may nonetheless be useful in applications that do not require
rigidity, e.g., tissue culture scaffolds and the like.
The mechanical property measurements of the POC-HA composites were
within the range of values reported for biodegradable polymers and
composites used or proposed for bone fixation devices [23].
Reported mechanical properties for polymers and composites have
included bending, compression, tensile strengths, and shear
strengths, whose values ranged from 40-412 MPa, 78-130 MPa, 0.6-290
MPa and 19-250 MPa, respectively. Reported values for bending,
compression, and tensile moduli ranged from 1.6-124.4 GPa, 4.8-8.0
GPa, and 0.01-29.9 GPa, respectively[12, 23]. The POC-HA composites
tested in this study had bending, compression, tensile, and shear
strengths that ranged from 33.9-41.4 MPa, 32-75 MPa, 6-10 MPa, and
23-28 MPa, respectively. Bending, compression, and tensile moduli
for POC-HA composites ranged from 0.275-0.502 GPa, 0.19-0.45 GPa,
and 0.02-0.34 GPa, respectively. Except for the bending and
compression moduli, the mechanical properties of POC-HA are
comparable to those of other biomaterials proposed for bone
fixation.
The mechanical properties of the POC-HA composites were increased
by increasing the HA component. It is also possible to modulate the
mechanical properties with the reaction conditions, i.e., reaction
temperature and time, and choice of diol for the polycondensation
reaction. Given the teachings of the present invention, the
composites of the invention can readily be adapted for in vivo use
in the intended application. The polymer-HA composites are expected
to integrate with bone. It is contemplated that the mass percent
and rate of degradation of the polymer component are parameters
that may influence the function and in vivo integration of the
composite. POC samples that were synthesized under the same
conditions as the synthesis of the POC-HA composites lost 46 wt. %
of their mass in 3 months. Further, POC when not combined with HA
has previously been shown to completely degrade within 6 months
when incubated in PBS at 37.degree. C. [15]. Without being bound by
any theory or mechanism of action, it is possible that the lower
degradation rates reported for the POC-HA composites may be due to
differing extents of the polycondensation reaction due to the lower
mass percentages of polymer and the presence of thermally
conductive HA particles. Both of those parameters are expected to
affect the degree of cross-linking relative to pure POC for the
same reaction temperature and time. Furthermore, it is also
possible for the POC to covalently react with OH groups on the HA
particles effectively crosslinking the POC-HA matrix [24-26]. HA
would also serve as a buffer to the acidic functional groups and
products generated from POC degradation, minimizing any
autohydrolytic effect on degradation. Further, the degradation of
the POC component can be significantly increased by "doping" with
glycerol or N-methyldiethanolamine (MDEA)[16].
In vivo, HA has been shown to induce the deposition of calcium
phosphate mineral on the surface of ceramic implants and bond to
bone [27, 28]. The capacity of POC and POC-HA to mineralize was
assessed in vitro using a modified simulated body fluid solution.
Based on the SEM and EDX analysis, POC-HA composites with 40-65 wt.
% HA in SBF successfully induced surface mineralization. The
mineralization process involved a nucleation phase and a growth
phase as evidenced by the complete coverage of the samples after 15
days of incubation in SBF [20]. However, POC was not conducive to
mineralization. The apatite or calcium phosphate mineral deposition
[29] may contribute to improved bone bonding in vivo and help fill
in any void volumes or pores left behind by degraded POC. Depending
on which bioceramic is chosen for a final application, most of the
mass of the screw is expected to be integrated (when the bioceramic
is an HA-type ceramic) or remodeled (when the bioceramic a TCP-type
ceramic) by bone tissue and the remaining POC should be totally
degraded within 2 years of implantation.
POC has been shown to be compatible (i.e. as per cell adhesion,
proliferation, and differentiation assays) with several cell types
including human and pig endothelial cells, human and pig smooth
muscle cells, bovine chondrocytes, and bovine fibroblasts [16, 30].
It was also shown to be biocompatible in vivo in a rat subcutaneous
implantation model [15]. In the present invention, the favorable
cell adhesion and spreading characteristics of POC and POC-HA
composites were confirmed in vitro with the use of primary human
osteoblasts. The cells adhered and formed a confluent monolayer on
all of the POC-HA composites evaluated (40, 50, 60, and 65 wt. %
HA) (see e.g. FIG. 5) providing evidence that the composites of the
present invention will be readily biocompatible in vivo. Such
determinations may further be corroborated using a bone defect
model.
From the above discussion, it is readily apparent that POC-HA
composites can be fabricated and molded into a variety of
orthopedic fixation devices. For example, POC-HA bone screws with
an HA content of 65 wt. % were successfully prepared. It will
readily be apparent that these compositions provide tremendous
advantages over the existing technologies. Advantages of such
POC-bioceramic composites include one or more of the following: a)
simple synthesis and in-situ crosslinking polymerization at
relatively mild temperatures while avoiding the use of exogenous
catalysts and toxic solvents, b) incorporation of a high percentage
of the bioceramic component, potentially enhancing
osteointegration, 3) a polymer component that should degrade
completely within two years rather than three to five years as in
the case with PLLA, and 4) decreased cost relative to the use of
poly(.alpha.-hydroxyl acids) such as PLLA. Moreover, the mechanical
properties of POC-HA composites can be adjusted with the percent of
HA in the composite and the material's surface supported
mineralization and osteoblast adhesion and proliferation. The
bioceramic particle size can be readily adjusted and its effects on
mechanical properties and in vivo bone integration characteristics
of the composite can be readily assessed.
In support of the above discussion, the following discussion
provides a further brief explanation of exemplary individual
components of the composites. Poly(diol citrates) are a family of
biodegradable and biocompatible elastomers that have shown
significant potential for soft tissue engineering, see e.g., U.S.
patent application see U.S. 60/721,687 and applications depending
therefrom. However, while those prior compositions are useful in
the production of matrices for tissue culture and implantable
tissue patches, those compositions are of insufficient rigidity to
serve in orthopedic indications. It is desirable to increase the
strength and stiffness of those composites in order to serve
orthopedic purposes. The methods of the present invention are
directed to strengthening such POC based polymers. Methods and
compositions for preparing POC are described in detail in
PCT/US2004/030631 and U.S. 60/721,687.
As noted herein, the POC is strengthened and stiffened for use in
orthopedic applications by incorporating bioceramics into the
POC-based elastomeric matrix. "Bioceramics" materials typically are
made of inorganic salts that include the ions of calcium and
phosphate, or in other examples include sulfate and carbonate.
Bioceramics fulfill a unique function as biomedical materials and
are used in a wide variety of applications in the human body.
Preferably, the bioceramics useful in the invention are
substantially non-toxic, biodegradable, bioerodable, and
bioresorbable. The terms "biodegradable" and "bioerodable" as used
herein similarly refer to a material property where biological,
biochemical, metabolic processes, and the like may effect the
erosion or degradation of the material over time. Such degradation
or erosion is due, at least in part, to contact with substances
found in the surrounding tissues, body fluids, and cells, or via
cellular action, enzymatic action, hydrolytic processes, and other
similar mechanisms in the body. The term "bioresorbable" as used
herein refers to materials that are used by, resorbed into, or are
otherwise eliminated from the body of the patient via existing
biochemical pathways and biological processes. For example, in
embodiments where the bioceramic comprises calcium phosphate,
bioresorbed calcium phosphate may be redeposited as bone mineral,
be otherwise reutilized within the body, or be excreted. It is
understood that some materials become bioresorbable following
biodegradation or bioerosion of their original state, as described
above. Preferably, the biocompatible material is such that it does
not elicit a substantial detrimental response in the host,
including but not limited to an immune reaction, such as an
inflammatory response, tissue necrosis, and the like that will have
a negative effect on the patient. In the event that such a negative
effect may be seen, preferably, the material may be treated with a
composition that allows the host to avoid such an adverse response
to the material.
Bioceramic materials typically are made of salts of alumina;
zirconia; calcium phosphates; silica based glasses or glass
ceramics; or pyrolytic carbons. The salts used to prepare the
bioceramics and the bioceramic matrices fabricated therefrom are
commercially available or are readily prepared via known
procedures. Bioceramics include calcium salts of carbonate,
sulfate, phosphate, and the like. Exemplary bioresorbable calcium
salts effective in the composition of this invention include
calcium carbonate, calcium sulfate, calcium sulfate hemihydrate,
also known as plaster of Paris, and certain porous or precipitated
forms of calcium phosphate, and the like. The porous bioceramic
matrix may also be fabricated from any number of natural bone
sources, such as autograft or allograft material, or synthetic
materials that are compositionally related to natural bone.
Calcium phosphate ceramics are in general prepared by sintering
more soluble calcium salts, for example Ca(OH).sub.2, CaCO.sub.3,
and CaHPO.sub.4, with a phosphorus-containing compound such as
P.sub.2O.sub.5. Such preparations of calcium phosphate ceramics are
known to those of skill in the art and have been described e.g., in
U.S. Pat. Nos. 3,787,900; 4,195,366; 4,322,398; 4,373,217 and
4,330,514 (each incorporated herein by reference). Exemplary
calcium phosphates for use in the invention include, but are not
limited to, calcium metaphosphate, dicalcium phosphate dihydrate,
calcium hydrogen phosphate, tetracalcium phosphates (TCPs),
heptacalcium decaphosphate, tricalcium phosphates, calcium
pyrophosphate dihydrate, crystalline hydroxyapatite, poorly
crystalline apatitic calcium phosphate, calcium pyrophosphate,
monetite, octacalcium phosphate, and amorphous calcium
phosphate.
Chemical formulae for calcium phosphate ceramics also are provided
in a range of crystalline morphologies, all of which may be used in
fabricating the bioceramic matrix, as described by U.S. Pat. Nos.
6,331,312 and 6,027,742. Such calcium phosphates have been
described as poorly-crystalline calcium phosphate (PCA) with an
apatitic structure. Other examples include tricalcium phosphate,
tetracalcium phosphate and other mixed-phase or polycrystalline
calcium phosphate materials reported in U.S. Pat. Nos. 4,880,610
and 5,053,312 to Constanz et al., the disclosures of which are
incorporated herein by reference.
Particularly preferred bioceramics for use in the present invention
include calcium phosphate apatites, such as hydroxyapatite (HA,
Ca.sub.10(PO.sub.4).sub.6(OH).sub.2) described by R. E. Luedemann
et al., Second World Congress on Biomaterials (SWCB), Washington,
D.C., 1984, p. 224, fluoroapatites, tricalciumphosphates (TCP),
such as Synthograft, dicalciumphosphates (DCP), and mixtures of HA
and TCP, as described by E. Gruendel et al., ECB, Bologna, Italy,
1986, Abstracts, p. 5, p. 32); mixed-metal salts such as magnesium
calcium phosphates, and beta-TCMP, as described by A. Ruggeri et
al., Europ. Congr. on Biomaterials (ECB), Bologna, Italy, 1986,
Abstracts, p. 86; aluminum oxide ceramics; bioglasses such as
SiO.sub.2--CaO--Na.sub.2O--P.sub.2O.sub.5, e.g. Bioglass 45S
(SiO.sub.2 45 wt %. CaO 24.5%, Na.sub.2O 24.5% and P.sub.2O.sub.5
6%) described by C. S. Kucheria et al., SWBC, Washington, D.C.,
1984, p. 214, and glass ceramics with apatites (MgO 4.6 wt %, CaO
44.9%, SiO.sub.2 34.2%, P.sub.2O.sub.5 16.3% and CaF 0.5%)
described by T. Kokubo et al., SWBC, Washington, D.C., 1984, p.
351; bioceramics incorporating organic ions, such as citrate, as
described in U.S. Pat. No. 5,149,368 to Liu et al.; and commercial
materials, such as Durapatite, Calcitite, Alveograf, and
Permagraft; the disclosures of which are incorporated herein by
reference.
In addition to the calcium-based bioceramics, it is contemplated
that bioactive glass compositions may also be used. Such
bioceramics include SiO.sub.2, Na.sub.2O, CaO, P.sub.2O.sub.5,
Al.sub.2O.sub.3, and CaF.sub.2. It is appreciated that the
above-described calcium salts may be used alone or may be mixed to
prepare the bioceramics described herein. Alumina and Zirconia are
known for their general chemical inertness and hardness. These
properties are exploited for implant purposes, where it is used as
an articulating surface in hip and knee joints. The ability of
these materials to be polished to a high surface finish makes them
ideal candidates for this wear application. Porous alumina has also
been used as a bone spacer, where sections of bone have had to be
removed due to disease. In this application, it acts as a scaffold
for bone ingrowth. Single crystal alumina or sapphire has also been
used.
Pyrolytic carbon is a bioceramic commonly used in artificial heart
valves and has been the most popular material for this application
for the last 30 years. Properties that make this material suitable
for this application include good strength, wear, resistance and
durability, and most importantly, thromboresistance, or the ability
to resist blood clotting. Pyrolytic carbon is also used for small
orthopaedic joints such as fingers and spinal inserts.
The POC-bioceramic materials may be prepared as porous or channeled
structures. It is understood that the nature and size of these
pores or channels may affect bioresorption. In certain aspects, the
pores of the structure are interconnected forming an open-cell
porous structure. It is understood that each of the foregoing
materials may possess differing bioresorption characteristics
obtainable in the treatment subject and such characteristics may be
advantageously chosen via routine experimentation for particular
variations of the processes and methods described herein. It is
also understood that both chemical composition and crystal
morphology may affect bioresorption rates. For example, bioceramics
fabricated from mixtures of calcium phosphate and calcium carbonate
or calcium phosphate and calcium sulfate typically undergo
resorption at higher rates than bioceramics fabricated from calcium
phosphate alone. Furthermore, highly crystalline bioceramics
typically undergo resorption at rates slower than poorly
crystalline or amorphous bioceramics.
The porous microstructure of the bioceramics may be achieved by
heat consolidation or sintering of bioceramic powders in
appropriate molds. The porous matrices may be macroporous or
microporous. Microporous matrices typically have pores in the range
from about 1 to about 100 microns in size, while macroporous
matrices typically have pores in the range from about 100 to about
1000 microns in size. In certain embodiments the pore size in a
given range is substantially uniform. The pores in the matrix
account for the void volume thereof. Such void volume may be from
about 30% to about 80%, and illustratively about 50% to about 70%
of the matrix volume. The pores are typically interconnecting, and
in some cases to a substantial degree. The pores may form an
open-cell configuration in some embodiments. In embodiments where
the void volume constitutes a substantial portion of the matrix
volume, the pores are typically close together. Illustratively,
adjacent pores are separated by less than 100 microns, and in other
embodiments separated by less than about the average of the
diameters of the adjacent pores.
In embodiments where the POC-bioceramic composites are used for the
preparation of bone-replacement materials, the pores may be
arranged in predetermined patterns that correspond to bone-healing
or bone-remodeling patterns, Haversian systems, and other
naturally-occurring patterns in bone. Commercially available
bioceramic matrices include e.g., Pro Osteon 200 and Pro Osteon 500
(hydroxyapatite bone-graft substitutes having interconnected porous
structures with pore sizes of 200 or 500 microns, similar to that
of cancerous bone) available from Interpore International, Irvine,
Calif.; Vitoss Blocks (calcium phosphate porous structure having
ca. 90% porosity, with pore sizes from 1 to 1000 microns in
diameter) available from Orthovita Inc., Malvern, Pa.; and
synthetic porous hydroxyapatite (made by a patented foam process
having controlled porosity and pore sizes) available from Hi-Por
Ceramics, United Kingdom. Such compositions may readily be used in
the present invention.
While the present invention is directed to production of components
for orthopedic devices, it is contemplated that in addition to the
bioceramic, the POC compositions also may be reinforced with a
second material. For example, the POC-HA composites also may
comprise a second biodegradable and biocompatible polymer. Two such
polymers are poly(L-lactic acid) (PLLA) and poly(lactic-co-glycolic
acid) (PLGA). These polymers are rigid and strong and have been
used in many tissue engineering applications. Furthermore, the rate
of degradation could be tailored to match that of the surrounding
elastomeric matrix. Poly(L-acetic acid) has a degradation time of
greater than two years while poly(glycolic acid) has a degradation
time of 1-2 months. By changing the ratio of lactic to glycolic
acid, the degradation rate could be varied from fast (1-2 months)
to slow (>2 years). For tissue engineering, the rate of
degradation of the polymer scaffold should match that of tissue
regrowth.
The reinforcing polymer may be a biodegradable polymer or a
non-biodegradable polymer but preferably is a biodegradable
polymer. Biodegradable polymers include, but are not limited to
collagen, elastin, hyaluronic acid and derivatives, sodium alginate
and derivatives, chitosan and derivatives gelatin, starch,
cellulose polymers (for example methylcellulose,
hydroxypropylcellulose, hydroxypropylmethylcellulose,
carboxymethylcellulose, cellulose acetate phthalate, cellulose
acetate succinate, hydroxypropylmethylcellulose phthalate), casein,
dextran and derivatives, polysaccharides, poly(caprolactone),
fibrinogen, poly(hydroxyl acids), poly(L-lactide) poly(D,L
lactide), poly(D,L-lactide-co-glycolide),
poly(L-lactide-co-glycolide), copolymers of lactic acid and
glycolic acid, copolymers of .epsilon.-caprolactone and lactide,
copolymers of glycolide and .epsilon.-caprolactone, copolymers of
lactide and 1,4-dioxane-2-one, polymers and copolymers that include
one or more of the residue units of the monomers D-lactide,
L-lactide, D,L-lactide, glycolide, .epsilon.-caprolactone,
trimethylene carbonate, 1,4-dioxane-2-one or 1,5-dioxepan-2-one,
poly(glycolide), poly(hydroxybutyrate), poly(alkylcarbonate) and
poly(orthoesters), polyesters, poly(hydroxyvaleric acid),
polydioxanone, poly(ethylene terephthalate), poly(malic acid),
poly(tartronic acid), polyanhydrides, polyphosphazenes, poly(amino
acids). The biodegradable polymers used herein may be copolymers of
the above polymers as well as blends and combinations of the above
polymers. (see generally, Illum, L., Davids, S. S. (eds.) "Polymers
in Controlled Drug Delivery" Wright, Bristol, 1987; Arshady, J.
Controlled Release 17:1-22, 1991; Pitt, Int. J. Phar. 59:173-196,
1990; Holland et al., J. Controlled Release 4:155-0180, 1986).
In particular preferred embodiments, the biodegradable or
resorbable polymer is one that is formed from one or more monomers
selected from the group consisting of lactide, glycolide,
.epsilon.-caprolactone, trimethylene carbonate, 1,4-dioxan-2-one,
1,5-dioxepan-2-one, 1,4-dioxepan-2-one, hydroxyvalerate, and
hydroxybutyrate. In one aspect, the polymer may include, for
example, a copolymer of a lactide and a glycolide. In another
aspect, the polymer includes a poly(caprolactone). In yet another
aspect, the polymer includes a poly(lactic acid),
poly(L-lactide)/poly(D,-L-Lactide) blends or copolymers of
L-lactide and D,L-lactide. In yet another aspect, the polymer
includes a copolymer of lactide and .epsilon.-caprolactone. In yet
another aspect, the polymer includes a polyester (e.g., a
poly(lactide-co-glycolide). The poly(lactide-co-glycolide) may have
a lactide:glycolide ratio ranges from about 20:80 to about 2:98, a
lactide:glycolide ratio of about 10:90, or a lactide:glycolide
ratio of about 5:95. In one aspect, the poly(lactide-co-glycolide)
is poly(L-lactide-co-glycolide; see e.g., U.S. Pat. No. 6,531,146
and U.S. application No. 2004/0137033.). Other examples of
biodegradable materials include polyglactin, and polyglycolic
acid.
Representative examples of non-biodegradable compositions include
ethylene-co-vinyl acetate copolymers, acrylic-based and
methacrylic-based polymers (e.g., poly(acrylic acid),
poly(methylacrylic acid), poly(methylmethacrylate),
poly(hydroxyethyl methacrylate), poly(alkylcynoacrylate),
poly(alkyl acrylates), poly(alkyl methacrylates)), polyolefins such
as poly(ethylene) or poly(propylene), polyamides (e.g., nylon 6,6),
poly(urethanes) (e.g., poly(ester urethanes), poly(ether
urethanes), poly(carbonate urethanes), poly(ester-urea)),
polyesters (e.g., PET, polybutyleneterephthalate, and
polyhexyleneterephthalate), olyethers (poly(ethylene oxide),
poly(propylene oxide), poly(ethylene oxide)-poly(propylene oxide)
copolymers, diblock and triblock copolymers, poly(tetramethylene
glycol)), silicone containing polymers and vinyl-based polymers
(polyvinylpyrrolidone, poly(vinyl alcohol), poly(vinyl acetate
phthalate), poly(styrene-co-isobutylene-co-styrene), fluorine
containing polymers (fluoropolymers) such as fluorinated ethylene
propylene (FEP) or polytetrafluoroethylene (e.g., expanded
PTFE).
The polymers may be combinations of biodegradable and
non-degradable polymers. Further examples of polymers that may be
used are either anionic (e.g., alginate, carrageenin, hyaluronic
acid, dextran sulfate, chondroitin sulfate, carboxymethyl dextran,
caboxymethyl cellulose and poly(acrylic acid)), or cationic (e.g.,
chitosan, poly-1-lysine, polyethylenimine, and poly(allyl amine))
(see generally, Dunn et al., J. Applied Polymer Sci. 50:353, 1993;
Cascone et al., J. Materials Sci.: Materials in Medicine 5:770,
1994; Shiraishi et al., Biol. Pharm. Bull. 16:1164, 1993;
Thacharodi and Rao, Int'l J. Pharm. 120:115, 1995; Miyazaki et al.,
Int'l J. Pharm. 118:257, 1995). Preferred polymers (including
copolymers and blends of these polymers) include
poly(ethylene-co-vinyl acetate), poly(carbonate urethanes),
poly(hydroxyl acids) (e.g., poly(D,L-lactic acid) oligomers and
polymers, poly(L-lactic acid) oligomers and polymers, poly(D-lactic
acid) oligomers and polymers, poly(glycolic acid), copolymers of
lactic acid and glycolic acid, copolymers of lactide and glycolide,
poly(caprolactone), copolymers of lactide or glycolide and
.epsilon.-caprolactone), poly(valerolactone), poly(anhydrides),
copolymers prepared from caprolactone and/or lactide and/or
glycolide and/or polyethylene glycol. Methods for making POC-PLLA
or PLGA or other like composites are described in U.S.
60/721,687.
The composites of the invention are load bearing and may bear loads
similar in magnitude to that borne by the tissue surrounding the
defect, such as a bone structure of similar dimensions, or a bone
structure consisting primarily of cortical bone. The structures
described herein may also possess mechanical properties similar to
that of natural bone, or in particular cortical bone. These
mechanical properties include, but are not limited to, tensile
strength, impact resistance, Young's modulus, compression strength,
sheer strength, stiffness, and the like. It is appreciated that
structures described herein possessing mechanical properties
similar to those exhibited by the tissue surrounding such implanted
structures may favorably influence the stress-shielding effect.
While it is appreciated that the above-described composites of the
invention will be fashioned into replacement bone or orthopedic
fixation devices, it is also contemplated that the devices may be
used as drug delivery system. Either the polymer, the bioceramic,
or both may include a biologically-active agent, either singly or
in combination, such that the composite structure or implant will
provide a delivery system for the agent at the site at which it is
implanted. Thus, the agent may advantageously be delivered to
adjacent tissues or tissues proximal to the implant site.
Biologically-active agents which may be used alone or in
combination in the implant precursor and implant include, for
example, a medicament, drug, or other suitable biologically-,
physiologically-, or pharmaceutically-active substance which is
capable of providing local or systemic biological, physiological,
or therapeutic effect in the body of the patient. The
biologically-active agent is capable of being released from the
solid implanted matrix into adjacent or surrounding tissue fluids
during biodegradation, bioerosion, or bioresorption of the fixation
device or artificial bone made from the composites of the
invention.
Incorporation of bioceramic composites of the invention supports
mineralization and osteoblast adhesion and proliferation, and can
potentially enhance osteointegration. In vivo bone integration
characteristics of the biodegradable elastomer-bioceramic
composites can be adjusted with different kinds and sizes of
bioceramic. Where the composites are used in facilitating bone
repair, the composites may advantageously be impregnated with an
"osteogenic agent" i.e., one which promotes, induces, stimulates,
generates, or otherwise effects the production of bone or the
repair of bone. The presence of an osteogenic agent in the site at
which the composite is placed may elicit an effect on the repair of
the defect in terms of shortening the time required to repair the
bone, by improving the overall quality of the repair, where such a
repair is improved over situations in which such osteogenic agents
are omitted, or may achieve contemporaneously both shortened repair
times and improved bone quality. It is appreciated that osteogenic
agents may effect bone production or repair by exploiting
endogenous systems, such as by the inhibition of bone
resorption.
Thus, osteogenic agents in the composites of the invention may be
used to effect repair of the bone by stabilizing the defect to
promote healing thereby increasing healing rate, producing a more
rapid new bone ingrowth, and improving overall repair of the bone.
The osteogenic agents may be synthetic molecules, drugs, or
pharmaceuticals involved in, or important to, bone biology,
including statins, such as lovastatin, simvastatin, atorvastatin,
and the like, fluprostenol, vitamin D, estrogen, a selective
estrogen receptor modifier, or a prostaglandin, such as PGE-2.
Growth factors or other proteins, peptides, receptor ligands,
peptide hormones, lipids, or carbohydrates involved in, or
important to, bone physiology may be used, including the bone
morphogenic or bone morphogenetic proteins (BMPs), such as BMP-2,
BMP-7, and BMP-9, chrysalin, osteogenic growth peptide (OGP), bone
cell stimulating factor (BCSF), KRX-167, NAP-52, gastric
decapeptide, parathyroid hormone (PTH), a fragment of parathyroid
hormone, osteopontin, osteocalcin, a fibroblast growth factor
(FGF), such as basic fibroblast growth factor (bFGF) and FGF-1,
osteoprotegerin ligand (OPGL), platelet-derived growth factor
(PDGF), an insulin-like growth factor (IGF), such as IGF-1 and
IGF-2, vascular endothelial growth factor (VEGF), transforming
growth factor (TGF), such as TGF-alpha and TGF-beta, epidermal
growth factor (EGF), growth and differentiation factor (GDF), such
as GDF-5, GDF-6, and GDF-7, thyroid-derived chondrocyte stimulation
factor (TDCSF), vitronectin, laminin, amelogenin, amelin, fragments
of enamel, or dentin extracts, bone sialoprotein, and analogs and
derivatives thereof.
In another embodiment the osteogenic agent is a cell or population
of cells involved in, or important to, bone biology, such as
pluripotent stem cells, autologous, allogenic, or xenogeneic
progenitor cells, chondrocytes, adipose-derived stem cells, bone
marrow cells, mesenchymal stem cells, homogenized or comminuted
tissue transplants, genetically transformed cells, and the like.
Bone powders, including demineralized bone powders and bone matrix,
may also be used. Combinations of such cell populations providing
the osteogenic agent are also contemplated herein.
The osteogenic agent may be present in the structure within the
range from about 0.1% to about 30% by weight, preferably in the
range from about 1% to 9% by weight.
Other agents also may be used in the composites. It is contemplated
that such additives may serve to reduce barriers to repair and thus
maximize the potential of the osteogenic agent. Preferably, such
agents are capable of preventing infection in the host, either
systemically or locally at the defect site, are contemplated as
illustrative useful additives. These additives include
anti-inflammatory agents, such as hydrocortisone, dexamethasone,
prednisone, and the like, NSAIDS, such as acetaminophen, salicylic
acid, ibuprofen, and the like, selective COX-2 enzyme inhibitors,
antibacterial agents, such as penicillin, erythromycin, polymyxin
B, viomycin, chloromycetin, streptomycins, cefazolin, ampicillin,
azactam, tobramycin, cephalosporins, bacitracin, tetracycline,
doxycycline, gentamycin, quinolines, neomycin, clindamycin,
kanamycin, metronidazole, and the like, antiparasitic agents such
as quinacrine, chloroquine, vidarabine, and the like, antifungal
agents such as nystatin, and the like, antiviricides, particularly
those effective against HIV and hepatitis, and antiviral agents
such as acyclovir, ribarivin, interferons, and the like. Systemic
analgesic agents such as salicylic acid, acetaminophen, ibuprofen,
naproxen, piroxicam, flurbiprofen, morphine, and the like, and
local anaesthetics such as cocaine, lidocaine, bupivacaine,
xylocalne, benzocaine, and the like, also can be used as additives
in the composites.
The composites of the present invention are used to fabricate
structures for use in orthopedic applications. Such structures
preferably are bone fixation devices, e.g., bone screws, pins,
plates and the like. Alternatively, the structures are those that
can be used for repairing bone voids, fractures, non-union
fractures, periodontal defects, maxillofacial defects, arthrodesis,
and the like. The composites may be fabricated into useful
structures using e.g., compression molding and machining methods
can be applied to fabricate any desirable shapes of fixation
devices and scaffolds for orthopaedic surgery and tissue
engineering. The mechanical properties and degradation of the
biodegradable elastomer bioceramic composites can be adjusted with
the percent and particle size (from micron to nanometer) of any
bioceramics (HA, TCP, OCP and bioglass) in the composite besides
ratio of diol and citric acid and post-polymerization conditions
such as temperature, vacuum, and polymerization time.
In addition, structures described herein may be used as
reinforcement of bone fractures, dental implants, bone implants,
bone prostheses and the like. It is appreciated that structures
described herein may also be generally used in conjunction with
other traditional fixation, immobilization, and prosthetic methods.
Fractures that may be treated by structures made from the
composites of the invention include fractures of the proximal
humerus, diaphyseal humerus, diaphyseal femur, trochanteric femur,
and trochanteric humerus. In addition, the structures may be used
in the repair of osteoporosis-induced fractures, including those
that involve a crushing-type injury, such as vertebral fractures,
and the like. In such fractures, the porous osteoporotic bone
collapses into itself typically causing a void or bone defect at
the site of the fracture, in order to achieve secure stabilization
of the fracture.
The various structures that can be prepared using the composites of
the invention may be fabricated by using methods known to those of
skill in the art. Typically, the bone fixation devices or other
structures for use in the orthopedic applications described herein
may be fabricated by compression molding. Compression molding
processes include transfer molding and squeeze-flow molding.
In exemplary embodiments, compression molding is used. A composite
of the invention, in a machined-block form, is placed on top of a
bioceramic matrix in a mold cavity. The mold is then heated to a
temperature at about or above the melting temperature of the
polymer. Minimal loading occurs during the heating step.
Pressurization of the mold is initiated once the molten polymer is
fluid enough for diffusion through the porous structure. In
addition, vacuum may be optionally applied during this process to
prevent degradation or hydrolysis of biocompatible polymer. It is
appreciated that applying a vacuum may also facilitate the
diffusion of polymer into the matrix. Exemplary molding of the
composites into bone screws is further described in the examples
herein below.
In other embodiments, transfer molding is used. A composite of the
invention, in a machined-block form, is preheated to a temperature
at about or above the melting temperature of the polymer and
subsequently transferred to a preheated mold cavity containing a
porous bioceramic matrix. Once the molten polymer is positioned,
squeeze molding is initiated by applying a load to a plunger,
thereby pressurizing the mold cavity.
In still other embodiments, flow molding is used. A porous
bioceramic matrix having a small-diameter open core is used. In
addition, the matrix has interconnected channels that are also
connected to the open core. The channels are arranged in a
substantially radial pattern when viewed in a given cross section
of the matrix. The porous matrix is placed in a mold cavity and the
POC material is disposed into the open core by either of the
above-described methods of compression molding or transfer molding.
In either case this process allows orientation of the polymer from
in the matrix. Such orientation may further reinforce and favorably
influence the mechanical properties of the structures described
herein.
Once the POC is disposed in the porous matrix by any of the methods
described herein, including compression molding, transfer molding,
squeeze-flow molding, and in-situ polymerization, the polymer may
be optionally crosslinked. Cross-linking may be accomplished by any
of the variety of known methods, including treatment with heat or
irradiation, such as X-ray radiation, gamma irradiation, electron
beam radiation, and the like.
It should be understood that the composites of the invention may be
provided to a practitioner as bulk material that may be shaped by
the medical practitioner on site. Alternatively, various
prefabricated shapes ready or near ready for implantation may be
produced from the composites. Such bulk material may in the form of
bars, blocks, billets, sheets, and the like. Such shapes include
plates, plugs, cubes, cylinders, pins, tubes, chutes, rods, screws,
including the screws described in U.S. Pat. No. 6,162,225 (bone
screw fabricated from allograft bone) the disclosure of which is
incorporated herein by reference, and the like. In addition, shapes
that tend to mimic the overall dimensions of the bone may be made
from the composites of the invention. Shapes that tend to mimic the
overall dimensions of the bone are particularly useful in the
repair of fractures at risk of non-union. Such bulk shapes or
particularly-dimensioned shapes may be obtained by employing mold
cavities possessing such dimensions. Alternatively, the
particularly-dimensioned shapes may be fabricated by machining the
bulk stock.
EXAMPLE 1
Biodegradable Elastomeric Polymers
The compositions of the invention are based on biodegradable
elastomeric polymers of poly(diol) citrate molecules. Such
molecules typically comprising a polyester network of citric acid
copolymerized with a linear aliphatic di-OH monomer in which the
number of carbon atoms ranges from 2 to 20. Polymer synthesis
conditions for the preparation of these molecules vary from mild
conditions, even at low temperature (less than 100.degree. C.) and
no vacuum, to tough conditions (high temperature and high vacuum)
according the requirements for the materials properties. By
changing the synthesis conditions (including, but not limited to,
post-polymerization temperature, time, vacuum, the initial monomer
molar ratio, and the di-OH monomer chain length) the mechanical
properties of the polymer can be modulated over a wide range. This
series of polymers exhibit a soft, tough, biodegradable,
hydrophilic properties and excellent biocompatibility in vitro.
The poly(diol)citrate polymers used herein have a general structure
of: (A-B-C).sub.n where A is a linear, aliphatic diol and C also is
a linear aliphatic diol. B is citric acid. The citric acid
co-polymers of the present invention are made up of multiples of
the above formula, as defined by the integer n, which may be any
integer greater than 1. It is contemplated that n may range from 1
to about 1000 or more. It is particularly contemplated that n may
be 1, 2, 3, 4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15, 16, 17, 18,
19, 20, 21, 22, 23, 24, 25, 26, 27, 28, 29, 30, 31, 32, 33, 34, 35,
36, 37, 38, 39, 40, 41, 42, 43, 44, 45, 46, 47, 48, 49, 50, or
more. In preferred embodiments, the compositions of poly(diol
citrates) that are used to prepare the implantable medical device
comprise a citric acid polyester having the generic formula
(A-B-C)n, wherein A and C could be any of the diols or any
combination of the diols; B could be citric acid, malic acid or
their combinations. The diols include linear or noon-linear
aliphatic diols, branched diol, cyclodiol, triol, heteroatom
containing diol (such as N-methyldiethanolamine, MDEA) and
macrodiol or their combinations. Any medical device coated with any
biodegradable elastomers (e.g., poly diol-citric acid,
polyurethanes, polycaprolactone and copolymers thereof) is
contemplated to be within the aspects of the present invention.
In preferred embodiments, the identity of "A" above is poly
1,10-decanediol and in another preferred embodiment the identity of
A is 1,8-octanediol. However, it should be understood that this is
merely an exemplary linear, aliphatic diol. Those of skill in the
art are aware of other aliphatic alcohols that will be useful in
polycondensation reactions to produce poly citric acid polymers.
Exemplary such aliphatic diols include any diols of between about 2
carbons and about 20 carbons. While the diols are preferably
aliphatic, linear, unsaturated diols, with the hydroxyl moiety
being present at the C.sub.1 and C.sub.x position (where x is the
terminal carbon of the diol), it is contemplated that the diol may
be an unsaturated diol in which the aliphatic chain contains one or
more double bonds. The preferred identity for "C" in one embodiment
is 1,8, octanediol, however as with moiety "A," "C" may be any
other aliphatic alcohols. While in specific embodiments, both A and
C are both the same diol, e.g., 1,8-octanediol, it should be
understood that A and C may have different carbon lengths. For
example, A may be 2, 3, 4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15,
16, 17, 18, 19, 20 or more carbons in length, and C may
independently be 2, 3, 4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15,
16, 17, 18, 19, 20 or more carbons in length. Exemplary methods for
the polycondensation of the citric acid with the linear diols are
provided in this Example. These materials are then used as starting
materials for the composites described in Example 2.
Synthesis of Poly(1,10-decanediol-co-citric acid) (PDC)
In a typical experiment, 19.212 g citric acid and 17.428 g
1,10-decanediol were added to a 250 ml three-neck round-bottom
flask, fitted with an inlet adapter and an outlet adapter. The
mixture was melted within 15 min by stirring at 160-165.degree. C.
in silicon oil bath, and then the temperature of the system was
lowered to 120.degree. C. The mixture was stirred for half an hour
at 120.degree. C. to get the pre-polymer. Nitrogen was vented
throughout the above procedures. The pre-polymer was
post-polymerized at 60.degree. C., 80.degree. C. or 120.degree. C.
with and without vacuum for predetermined time from one day to 3
weeks depending on the temperature to achieve the
Poly(1,10-decanediol-co-citric acid). Nitrogen was introduced into
the reaction system before the polymer was taken out from the
reaction system.
Preparation of Poly(1,8-Octanediol-co-citric acid) (POC)
In a typical experiment, 19.212 g citric acid and 14.623 g
1,8-octanediol were added to a 250 mL three-neck round-bottom
flask, fitted with an inlet adapter and an outlet adapter. The
mixture was melted within 15 min by stirring at 160-165.degree. C.
in silicon oil bath, and then the temperature of the system was
lowered to 140.degree. C. The mixture was stirred for another 1 hr
at 140.degree. C. to get the pre-polymer. Nitrogen was vented
throughout the above procedures. The pre-polymer was
post-polymerized at 60.degree. C., 80.degree. C. or 120.degree. C.
with and without vacuum for predetermined time (from one day to 3
weeks depending on the temperature, with the lower temperatures
requiring longer times) to achieve the
Poly(1,8-octanediol-co-citric acid). Nitrogen was introduced into
the reaction system before the polymer was taken out from the
reaction system.
Porous scaffolds of POC (tubular and flat sheets) were prepared via
a salt leaching technique as follows: POC pre-polymer was dissolved
into dioxane to form 25 wt % solution, and then the sieved salt
(90-120 microns) was added into pre-polymer solution to serve as a
porogen. The resulting slurry was cast into a
poly(tetrafluoroethylene) (PTFE) mold (square and tubular shape).
After solvent evaporation for 72 h, the mold was transferred into a
vacuum oven for post-polymerization. The salt in the resulting
composite was leached out by successive incubations in water
(produced by Milli-Q water purification system every 12 h for a
total 96 h. The resulting porous scaffold was air-dried for 24 hr
and then vacuum dried for another 24 hrs. The resulting scaffold
was stored in a desiccator under vacuum before use. Porous
scaffolds are typically preferred when cells are expected to
migrate through a 3-dimensional space in order to create a tissue
slice. Solid films would be used when a homogenous surface or
substrate for cell growth is required such as an endothelial cell
monolayer within the lumen of a vascular graft.
Using similar techniques porous scaffold of PDC or other
poly(diol)citrates can be prepared. In other embodiments, biphasic
scaffolds can be prepared. Biphasic scaffolds consist of an outside
porous phase and an inside non-porous phase as depicted in the
schematic drawing shown in FIG. 15 of PCT PCT/US2004/030631,
incorporated herein by reference. The non-porous phase is expected
to provide a continuous surface for EC adhesion and spreading,
mechanical strength, and elasticity to the scaffold. The porous
phase will facilitate the 3-D growth of smooth muscle cells.
Biphasic scaffolds were fabricated via following procedures.
Briefly, glass rods (.about.3 mm diameter) were coated with the
pre-polymer solution and air dried to allow for solvent
evaporation. Wall thickness of the tubes can be controlled by the
number of coatings and the percent pre-polymer in the solution. The
pre-coated pre-polymer was partially post-polymerized under
60.degree. C. for 24 hr; the pre-polymer-coated glass rod is then
inserted concentrically in a tubular mold that contains a
salt/pre-polymer slurry. The pre-polymer/outer-mold/glass rod
system is then placed in an oven for further post-polymerization.
After salt-leaching, the biphasic scaffold was then de-molded from
the glass rod and freeze dried. The resulting biphasic scaffold was
stored in a desiccator before use. The same materials or different
materials from the above family of elastomers can be utilized for
both phases of the scaffold. Other biomedical materials widely used
in current research and clinical application such as polylactide
(PLA), polycaptrolactone (PCL), poly(lactide-co-glycolide) (PLGA)
may also be utilized for this novel scaffold design.
The thickness, degradation, and mechanical properties of the inside
non-porous phase can be well controlled by choosing various
pre-polymers of this family of elastomers, pre-polymer
concentration, coating times and post-polymerization conditions
(burst pressure can be as high as 2800 mmHg). The degradable porous
phase and non-porous phases are integrated since they are formed
in-situ via post-polymerization. Cell culture experiments confirm
that both HAEC and HASMC can attach and grow well in biphasic
scaffolds. The results suggest that a biphasic scaffold design
based on poly(diol citrate) is a viable strategy towards the
engineering of small diameter blood vessels.
Synthesis of Poly(1,6-hexanediol-co-citric acid) (PHC)
In a typical experiment, 19.212 g citric acid and 11.817 g
1,6-hexanediol were added to a 250 ml three-neck round-bottom
flask, fitted with an inlet adapter and an outlet adapter. The
mixture was melted within 15 min by stirring at 160-165.degree. C.
in a silicon oil bath, and then the temperature of the system was
lowered to 120.degree. C. The mixture was stirred for half an hour
at 120.degree. C. to get the pre-polymer. Nitrogen was vented
throughout the above procedures. The pre-polymer was
post-polymerized at 60.degree. C., 80.degree. C. or 120.degree. C.
with and without vacuum for a predetermined time from one day to 3
weeks, depending on the temperature, to achieve the
Poly(1,6-hexanediol-co-citric acid). Nitrogen was introduced into
the reaction system before the polymer was taken out from the
reaction system.
Synthesis of Poly(1,12-dodecanediol-co-citric acid) PDDC
In a typical experiment, 19.212 g citric acid and 20.234 g
1,12-dodecanediol were added to a 250 ml three-neck round-bottom
flask, fitted with an inlet adapter and an outlet adapter. The
mixture was melted within 15 min by stirring at 160-165.degree. C.
in silicon oil bath, and then the temperature of the system was
lowered to 120.degree. C. The mixture was stirred for half an hour
at 120.degree. C. to get the pre-polymer. Nitrogen was vented
throughout the above procedures. The pre-polymer was
post-polymerized at 60.degree. C., 80.degree. C. or 120.degree. C.
with and without vacuum for predetermined time from one day to 3
weeks depending on the temperature to achieve the
Poly(1,12-dodecanediol-co-citric acid). Nitrogen was introduced
into the reaction system before the polymer was taken out from the
reaction system.
Synthesis of Poly(1,8-octanediol-co-citric acid-co-glycerol)
In a typical experiment (Poly(1,8-octanediol-co-citric acid-co-1%
glycerol), 23.0544 g citric acid, 16.5154 g 1,8-octanediol and
0.2167 g glycerol were added to a 250 ml three-neck round-bottom
flask, fitted with an inlet adapter and an outlet adapter. The
mixture was melted within 15 min by stirring at 160-165.degree. C.
in silicon oil bath, and then the temperature of the system was
lowered to 120.degree. C. The mixture was stirred for another hour
at 140.degree. C. to get the pre-polymer. Nitrogen was vented
throughout the above procedures. The pre-polymer was
post-polymerized at 60.degree. C., 80.degree. C. or 120.degree. C.
with and without vacuum for predetermined time from one day to 3
weeks depending on the temperature to achieve the
Poly(1,8-octanediol-co-citric acid-co-1% glycerol). Nitrogen was
introduced into the reaction system before the polymer was taken
out from the reaction system.
Synthesis of Poly(1,8-octanediol-citric acid-co-polyethylene
oxide)
In a typical experiment, 38.424 g citric acid, 14.623 g
1,8-octanediol and 40 g polyethylene oxide with molecular weight
400 (PEO400) (100 g PEO1000 and 200 g PEO2000 respectively) (molar
ratio: citric acid/1,8-octanediol/PEO400=1/0.5/0.5) were added to a
250 ml or 500 ml three-neck round-bottom flask, fitted with an
inlet adapter and an outlet adapter. The mixture was melted within
15 min by stirring at 160-165.degree. C. in silicon oil bath, and
then the temperature of the system was lowered to 135.degree. C.
The mixture was stirred for 2 hours at 135.degree. C. to get the
pre-polymer. Nitrogen was vented throughout the above procedures.
The pre-polymer was post-polymerized at 120.degree. C. under vacuum
for predetermined time from one day to 3 days to achieve the
Poly(1,8-octanediol-citric acid-co-polyethylene oxide). Nitrogen
was introduced into the reaction system before the polymer was
taken out from the reaction system. The molar ratios can be altered
to achieve a series of polymers with different properties.
Synthesis of Poly(1,12-dodecanediol-citric acid-co-polyethylene
oxide)
In a typical experiment, 38.424 g citric acid, 20.234 g
1,12-dodecanediol and 40 g polyethylene oxide with molecular weight
400 (PEO400) (100 g PEO1000 and 200 g PEO2000 respectively) (molar
ratio: citric acid/1,8-octanediol/PEO400=1/0.5/0.5) were added to a
250 ml or 500 ml three-neck round-bottom flask, fitted with an
inlet adapter and an outlet adapter. The mixture was melted within
15 min by stirring at 160-165.degree. C. in silicon oil bath, and
then the temperature of the system was lowered to 120.degree. C.
The mixture was stirred for half an hour at 120.degree. C. to get
the pre-polymer. Nitrogen was vented throughout the above
procedures. The pre-polymer was post-polymerized at 120.degree. C.
under vacuum for predetermined time from one day to 3 days to
achieve the Poly(1,12-dodecanediol-citric acid-co-polyethylene
oxide). Nitrogen was introduced into the reaction system before the
polymer was taken out from the reaction system. The molar ratios
can be altered to achieve a series of polymers with different
properties.
Synthesis of Poly(1,8-octanediol-citric
acid-co-N-methyldiethanoamine) POCM
In a typical experiment, 38.424 g citric acid, 26.321 g
1,8-octanediol and 2.3832 g N-methyldiethanoamine (MDEA) (molar
ratio: citric acid/1,8-octanediol/MDEA=1/0.90/0.10) were added to a
250 ml or 500 ml three-neck round-bottom flask, fitted with an
inlet adapter and an outlet adapter. The mixture was melted within
15 min by stirring at 160-165.degree. C. in silicon oil bath, and
then the temperature of the system was lowered to 13520.degree. C.
The mixture was stirred for half an hour at 120.degree. C. to get
the pre-polymer. Nitrogen was vented throughout the above
procedures. The pre-polymer was post-polymerized at 80.degree. C.
for 6 hours, 120.degree. C. for 4 hours without vacuum and then
120.degree. C. for 14 hours under vacuum to achieve the
Poly(1,8-octanediol-citric acid-co-N-methyldiethanoamine). Nitrogen
was introduced into the reaction system before the polymer was
taken out from the reaction system. The molar ratios can be altered
to citric acid/1,8-octanediol/MDEA=1/0.95/0.05.
Synthesis of Poly(1,12-dodecanediol-citric
acid-co-N-methyldiethanoamine) PDDCM
In a typical experiment, 38.424 g citric acid, 36.421 g
1,12-dodecanediol and 2.3832 g N-methyldiethanoamine (MDEA) (molar
ratio: citric acid/1,8-octanediol/MDEA=1/0.90/0.10) were added to a
250 ml or 500 ml three-neck round-bottom flask, fitted with an
inlet adapter and an outlet adapter. The mixture was melted within
15 min by stirring at 160-165.degree. C. in a silicon oil bath, and
then the temperature of the system was lowered to 120.degree. C.
The mixture was stirred for half an hour at 120.degree. C. to get
the pre-polymer. Nitrogen was vented throughout the above
procedures. The pre-polymer was post-polymerized at 80.degree. C.
for 6 hours, 120.degree. C. for 4 hours without vacuum and then
120.degree. C. for 14 hours under vacuum to achieve the
Poly(1,12-dodecanediol-citric acid-co-N-methyldiethanoamine).
Nitrogen was introduced into the reaction system before the polymer
was taken out from the reaction system. The molar ratios can be
altered to citric acid/1,12-dodecanediol/MDEA=1/0.95/0.05.
EXAMPLE 2
Materials and Methods Used in Preparing and Characterizing
Biodegradable Elastomeric Composites Made from POC and
Bioceramics
Example 1 describes the production of PDC as well as a number of
other poly(diol)citrate polymers. In the present Example, there are
provided teachings of how to further strengthen and stiffen is the
poly(diol)citrate polymers by incorporating ceramics into the
elastomeric polymer matrix.
Materials and Methods
Materials:
Hydroxyapatite [Mw: 502.32, Assay >90 (as Ca3 (PO4)2),
0.5%>75 um, 1.4% between 45-75 um, 98.1%<45 um] was purchased
from Fluka (St. Louis, Mo., USA). 1,8-octanediol (98%) and citric
acid (99.5%) were purchased from Sigma-Aldrich (St. Louis, Mo.,
USA). These materials were used as received. PTFE tubes were
purchased from McMaster-CARR, Chicago, USA.
Sample Preparation:
POC pre-polymer was synthesized according to published methods
[15]. Briefly, 0.05 mol of 1.8-octandiol and 0.05 mol of citric
acid were added to a 100 ml round bottom flask and exposed to a
constant flow of nitrogen gas. The mixture was melted under
vigorous stirring at 160-165.degree. C. Following melting, the
mixture was polymerized at 140.degree. C. for 1 hr to create a POC
pre-polymer. The POC pre-polymer was mixed with various amounts of
HA particles to obtain composites of 40, 50, 60, and 65 wt. % HA by
mass. Briefly, POC pre-polymer was mixed with the desired amount of
HA powder and placed in PTFE dishes that were pre-warmed to
80.degree. C. The POC-HA mixture was stirred until it became
clay-like, a process that generally took 5-10 hrs depending on the
HA content. The POC-HA mass was then inserted into PTFE tubes to
make rods or into other PTFE molds designed to meet the dimensional
requirements for sample mechanical testing protocols or in situ
formation of bone screws. The POC-HA in the mold was then
post-polymerized at 80.degree. C. for 3 days followed by
120.degree. C. under 2 Pa vacuum for 1 day.
Characterization of the mechanical properties of POC-HA composites:
The following mechanical properties were measured using a Sintech
mechanical tester model 20/G (Triangle Park, N.C. owned by MTS
now): 1) bending strength (Sb) and modulus (Eb) according to
Japanese industrial standard (JIS) K7203, 2) compression strength
(Sc) and modulus (Ec) according to JIS K7208, 3) tensile strength
(St) and modulus (Et) according to JIS K7113, 4) shear strength
(Ss)[17] and 5) torsional strength (Ts) [18].
All rods used for the mechanical tests were polished with sandpaper
before measurement. For all mechanical tests, at least 6 samples
were tested and the mean values and standard deviations (SD) were
calculated. The density of POC and POC-HA composites was measured
using the Archimedes principle as previously described [19].
Characterization of Morphology of POC-HA Composites:
SEM was used for observation of morphology. All POC-HA composites
for morphology were the cross section of rods obtained by
compression method.
Characterization of the In Vitro Degradation of POC-HA
Composites:
The degradation of POC-HA composite samples (10 mm diameter.times.2
mm thick) with HA percentages of 40, 50, 60, and 65 wt. % was
assessed in vitro in PBS, pH 7.4, at 37.degree. C. for up to 30
weeks under static conditions. Within the POC-HA composite, only
the POC is expected to degrade when incubated in aqueous solution.
For comparison purposes, the degradation of POC samples synthesized
under the same conditions as the composites was also assessed. PBS
was changed as necessary to ensure that the pH did not drop below
7. Prior to weighing, samples were rinsed with deionized water and
dried. Mass loss was calculated by comparing the initial mass (Wo)
with the mass measured at a given time point (Wt), as shown in
Equation 1. The results are presented as means.+-.standard
deviation (n=4).
.times..times..function..times. ##EQU00001##
Mineralization of POC-HA composites: Surface mineralization of
POC-HA composites was assessed in vitro using modified simulated
body fluid (SBF)[20]. The SBF consisted of (mmol): Na.sup.+
(142.0), K+(4.0), Mg.sup.2+ (1.5), Ca.sup.2+ (5.0), Cl.sup.- (147),
HCO.sub.3.sup.- (4.2), HPO.sub.4.sup.2- (2.0) and SO.sub.4.sup.2-
(0.5) with the pH adjusted to 7.2 using
tris(hydroxymethyl)aminomethane [21]. Briefly, discs (10 mm
diameter.times.2 mm thick) of POC-HA composites with HA fraction of
40, 50, 60 and 65 wt. % were immersed in 10 ml of the SBF at
37.degree. C. for up to 15 days. Fresh SBF was added every other
day to maintain the ionic concentration and pH during
mineralization. The morphology of deposited calcium phosphate
crystals was observed via scanning electron microscopy (SEM)
(Hitachi 3500 N, EPIC, Northwestern University). The stoichiometric
Ca/P molar ratio was analyzed by energy dispersive X-ray (EDX).
Evaluation of the cell compatibility of POC-HA composites: POC-HA
discs (7.0 mm in diameter.times.2 mm thick, with 40, 50, 60, and 65
wt. % HA) were sterilized by incubation in 70% ethanol for 30
minutes, washing with sterile PBS (pH 7.4), and UV exposure for 30
minutes. After sterilization, samples were washed several times
with cell culture media prior to placement in the wells of a
48-well tissue culture plate. A 40 .mu.l volume of a suspension of
human osteoblast cells (HOB) (Cambrex, Pittsburgh, Pa.)
(3.times.10.sup.5 cells mL.sup.-1) was added to each well and
incubated in osteoblast growth medium (OBM and OGM SingleQuots from
Cambrex) at 37.degree. C. in humidified air and 5% CO2 for up to 14
days. The culture medium was changed every three days. Samples were
fixed with 2.5% gluteraldehyde in PBS for 24 h at 4.degree. C. The
morphology of the cells on the composite samples was observed via
SEM.
Statistical Methods:
Data are expressed as means.+-.standard deviation. The statistical
significance was calculated using two-tail Student's t-test and
analysis of variance (ANOVA) and post-hoc analysis using one-way
analysis of variation (ANOVA): Newman-Keuls Multiple Comparison
Test. P<0.05 was considered as significant differences.
EXAMPLE 3
Preparation and Analysis of Bone Screws Made from Composites of POC
and Bioceramics
POC-HA composites having HA fraction of 40 wt. %, 50 wt. %, 60 wt.
%, 65 wt. % HA were investigated for bone screws. Using compression
molding method, POC-HA composites with HA from 40-65 wt. % were
compressed into molds of rods and screws. The POC-HA rods obtained
were strong enough to be further machined into screws. The POC-HA
screws from compression molding and machining methods were shown in
FIG. 1.
Characterization of the Mechanical Properties of Composites.
The mechanical property measurements (bending, compression, shear,
tension and torsion) are summarized in Table 1.
TABLE-US-00001 TABLE 1 Effect of HA fraction on mechanical
properties of POC-HA composites POC- HA .rho. Sb Eb Sc Ec St Et Ss
Ts Composite (Wt %) (g/cm.sup.3) (MPa) (MPa) (MPa) (MPa) (MPa)
(MPa) (MPa) (N.m) HA 40 40/60 1.609 33.9 275 32.0 189 7.8 21.4 23.3
22.9 (.+-.0.016) (.+-.5.7) (.+-.80) (.+-.13.0) (.+-.21) (.+-.0.5)
(.+-.1.8) (- .+-.1.6) (.+-.1.6) HA 50 50/50 1.653 37.7 323 64.0 264
7.1 30.2 25.1 24.2 (.+-.0.014) (.+-.4.6) (.+-.63) (.+-.9.4)
(.+-.14) (.+-.0.3) (.+-.2.2) (.- +-.1.7) (.+-.2.0) HA 60 60/40
1.734 34.7 314 52.6 297 6.4 85.4 25.9 21.4 (.+-.0.061) (.+-.3.0)
(.+-.54) (.+-.11.5) (.+-.41) (.+-.2.0) (.+-.8.8) (- .+-.1.6)
(.+-.1.9) HA 65 65/35 1.885 41.4 502 74.6 449 9.7 334.8 27.7 27:3
(.+-.0.072) (.+-.3.1) (.+-.40) (.+-.9.0) (.+-.27) (.+-.2.3)
(.+-.73.5) (- .+-.2.4) (.+-.4.9) Note: .rho.: density; Sb: bending
strength. Eb: bending modulus, Ss: shear strength. Sc: compression
strength. Ec: compression modulus, St: tensile strength
(rectangular specimens), Et: tensile modulus, Ts: torsional
strength
(1) Bending strength (Sb) and modulus (Eb): With increasing HA
fraction from 40 wt. % to 65 wt. %, bending strength (Sb) increased
reaching about 41 MPa at 65 wt. % which was considered as the
highest one among the composites, and composites showed almost the
similar Sb at 40, 50, and 60 wt. %. The bending modulus showed the
similar increasing tendencies as bending strength with increasing
HA fraction from 40-65 wt. %. The highest modulus is 502 MPa at 65
wt. %.
(2) Compression strength (Sc) and modulus (Ec): Compression
strength (Sc) greatly increased from 32 to 75 MPa in proportion to
HA fraction from 40 to 65 wt. % except for 50 wt. %. Compression
modulus at 65 wt. % HA is the highest and the values for composites
with 40-60 wt. % HA are similar to each other. It seems likely that
it originated from the HA particle aggregating in the
composites.
(3) Shear strength (Ss): The shear strength (Ss) at 65 wt. % is
significantly higher than that at 40 wt. %, and there is no
significant difference at 40-60 wt. %.
(4) Tensile strength (St) and modulus (Et): Compared with POC,
tensile strength improved greatly for composites with fraction from
40-65 wt. %, and the composite at 65 wt. % has higher tensile
strength than those at 50 wt. % and 60 wt. %. Modulus increased
with increasing HA concentration from 50-65 wt. % and reached 335
MPa at 65 wt. %, the highest one among composites.
(5) Torsional strength (Tt): Torsional strength (Tt) did not change
significantly with increasing HA fraction from 50-60 wt. %.
Torsional strength at 65 wt. % HA is the highest.
In summary, the POC-HA composites with 65 wt. % HA had the highest
bending strength and modulus, compression strength and modulus,
tensile strength and modulus, shear strength and torsional
strength. Interestingly, for HA percent of 40-60 wt. % there was no
statistically significant difference for most of the mechanical
properties. The mechanical properties of all POC-HA composites
tested were significantly increased relative to POC samples. The
density of samples increased with increasing percentage of HA.
EXAMPLE 4
The Morphology and Other Properties of POC-HA Composites
FIG. 2 showed that for the POC composite with 40 wt % HA, HA as
nodules dispersed in POC-HA. However, there were continuous flakes
of HA covering on the surface for the POC composite with 65 wt. %
HA, indicating higher amount of HA in this composite.
In Vitro Degradation of POC-HA Composites
The mass loss over time profiles for POC-HA composites incubated in
PBS at 37.degree. C. are shown in FIG. 3. Mass loss is due to the
aqueous hydrolytic degradation of POC within the composite. The
degradation of the POC-HA composite scaffolds with HA percentages
of 40, 50, and 60 was very similar at all time points with a total
mass loss of approximately 12 wt. % at 20 weeks. POC-HA composite
with 65 wt. % HA had the slowest degradation rate with a mass loss
of approximately 8.4 wt. % at 20 weeks. For comparison purposes,
POC samples lost 46 wt. % of their mass by 12 weeks.
Mineralization of POC and POC-HA Composites in SBF
SEM showed that mineralization was not observed on the surface of
POC incubated in SBF for 3 days through 15 days as shown in FIGS.
4a and b. However, mineral nodules began to form and aggregated on
the surface of composites throughout the 15 days of incubation in
SBF at 37.degree. C. (FIG. 4 c-f). Mineral nodules merged into a
continuous covering on most of the sample's surface at 15 days. The
composition of the mineral, in terms of the molar ratio of Ca/P,
was confirmed by the EDX to be 1.5-1.7. Deposition of a large
number of calcium phosphates was remarkable on the surface of
composites due to exposure of HA. The osteoconductivity of POC-HA
composites would be largely predicted.
Evaluation of the Cell Compatibility of POC-HA Composites
In order to investigate whether HOB can attach and proliferate on
the POC-HA composites, the morphology of cells seeded on the
surface of POC and POC-HA composites with 40 wt. % and 65 wt. % HA
respectively was observed by light microscopy (LM) and SEM as shown
in FIG. 5. LM showed that HOB cells can attach, spread out and
proliferate on the surface of POC cultured from 3 days through 14
days (FIG. 5a-b). SEM in FIGS. 5c and e showed that for both POC-HA
composites cultured for 3 days, the cells attached and spread out
well on the surface of both composites, and layers of the cells on
the surface of 65 wt. % POC-HA composite were observed. It revealed
a good covering on both composites, and better cell proliferation
on 65 wt. % HA composites. While culturing up to 14 days, both
surfaces were almost completely covered by layers of cells (FIG. 5
d and f). Moreover, the broken cell layers were observed on the
surfaces of both materials in some area due to the thickness of the
cell layers. However, HOB cells still remain round aggregating on
the POC surface even culturing for 14 days.
EXAMPLE 5
Further Measurements of Mechanical Properties
(1) Bending strength (Sb) and modulus (Eb). (Japanese Industrial
Standard (JIS) K7203 measurements were determined using a three
point bending test using rods with a range of diameter from 5.0 to
6.5 mm and a length of 30 mm). Sb=8FmaxL/.pi.D.sub.3
Eb=4L.sub.3/3.pi.D.sub.4(E/Y) where Fmax=maximum force (N);
L=support span (mm); D=diameter (mm); and E/Y=gradient of linear
portion of stress-strain curve (N/mm).
(2) Compression strength (Sc) and modulus (Ec) (Japanese Industrial
Standard (JIS) K7208 was measured using rods with a range of
diameter from 5.0 to 6.5 mm and a length of 15-30 mm. Sc=Fmax/A
Ec=E/Y where Fmax=maximum force (N); A=compressed area (mm2);
E/Y=gradient of linear portion of stress-strain curve (N/mm2).
(3) Shear strength (Ss) (Measured by Suuronen's method [20] at a
testing speed of 10 mm min.sup.-1) using rods with a range of
diameter from 5.0 to 6.5 mm and a length of 20 mm. Ss=Fmax/2A where
Fmax=maximum force (N); A=loading area (mm2).
(4) Tensile strength (St) and modulus (Et). (Japanese Industrial
Standard (JIS) K7113 using dog bone shaped samples (26 mm.times.4
mm.times.1.6 mm) at testing speed of 10 mm/min St=Fmax/A Et=E/Y
Where Fmax=maximum force (N); A=transversal cross-sectional area
(mm2); E/Y=gradient of linear portion of stress-strain curve
(N/mm).
(5) Torsional strength (Ts) The test rod was installed to Sintech
20/G Materials Testing Machine, using the dumbbell shaped samples
having an average diameter of 4.7 mm and length of 16.5 mm. The
rotating wheel was turned at a rate of 0.4 rev/min by means of a
chain attached to the load cell. The load pulling the chain was
recorded and used for the calculation of the torque strength
according to the equation Ts=16FmaxR/(.pi.D3) where Ts=torsion
strength (MPa), Fmax=load at fracture (N), R=radius of the rotating
wheel=38 mm and D=diameter of the rod (mm).
TABLE-US-00002 TABLE 2 Results of ANOVA on Mechanical Properties
P-values Sb Eb Sc Ec St Et Ss Ts Number (MPa) (MPa) (MPa) (MPa)
(MPa) (MPa) (MPa) (N.m) HA40 N S N S P<0.001 N S N S N S N S N S
vs HA50 HA40 N S N S P<0.05 N S N S P<0.001 N S N S vs HA60
HA40 P<0.05 P<0.001 P<0.001 P<0.01 N S P<0.001
P<0.01 P<0.05 vs HA65 HA50 N S N S P<0.05 N S N S
P<0.001 N S N S vs HA60 HA50 N S P<0.001 N S N S P<0.05
P<0.001 N.S N.S vs HA65 HA60 P<0.05 P<0.001 P<0.01
P<0.05 P<0.01 P<0.001 N S P<0.05 vs HA65 Note: Sb:
bending strength. Eb: bending modulus, Ss: shear strength. Sc:
compression strength. Ec: compression modulus, St: tensile strength
(rectangular specimens), Et: tensile modulus, Ts: torsional
strength. Tensile strength POC vs HA (40-65%): P<0.001; Tensile
modulus POC vs HA (60-65%): P<0.001; and POC vs HA (40-50%): N
S
TABLE-US-00003 TABLE 3 Results of ANOVA on degradation between
POC-HA (40-65 wt. %) composites respectively in 2 and 20 weeks.
P-values Time HA40 vs HA40 vs HA50 vs HA40 vs HA50 vs HA60 vs
(week) HA50 HA60 HA60 HA65 HA65 HA65 2 N.S N.S N.S P < 0.01 P
< 0.001 P < 0.001 20 N.S N.S N.S P < 0.05 P < 0.05 P
< 0.01
TABLE-US-00004 TABLE 4 Results of ANOVA on degradation of POC-HA
(40-65 wt. %) composites from 2 to 20 weeks. Time P-values (week)
HA40 HA50 HA60 HA65 2 vs 6 P < 0.001 P < 0.001 N.S N.S 2 vs
12 P < 0.001 P < 0.001 N.S N.S 2 vs 20 P < 0.001 P <
0.001 N.S P < 0.01 6 vs 12 P < 0.05 P < 0.01 N.S N.S 6 vs
20 P < 0.001 P < 0.001 N.S P < 0.01 12 vs 20 P < 0.01 P
< 0.05 P < 0.05 P < 0.05
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